Tissue, such as a benign or malignant tumor or blood clot within a patient's skull or other body region, may be treated invasively by surgically removing the tissue or non-invasively by using, for example, thermal ablation. Both approaches may effectively treat certain localized conditions within the body, but involve delicate procedures to avoid destroying or damaging otherwise healthy tissue. Unless the healthy tissue can be spared or its destruction is unlikely to adversely affect physiological function, surgery may not be appropriate for conditions in which diseased tissue is integrated within healthy tissue.
Thermal ablation, as may be accomplished using focused ultrasound, has particular appeal for treating diseased tissue surrounded by or neighboring healthy tissue or organs because the effects of ultrasound energy can be confined to a well-defined target region. Ultrasonic energy may be focused to a zone having a cross-section of only a few millimeters due to relatively short wavelengths (e.g., as small as 1.5 millimeters (mm) in cross-section at one Megahertz (1 MHz)). Moreover, because acoustic energy generally penetrates well through soft tissues, intervening anatomy often does not impose an obstacle to defining a desired focal zone. Thus, ultrasonic energy may be focused at a small target in order to ablate diseased tissue without significantly damaging surrounding healthy tissue.
An ultrasound focusing system generally utilizes an acoustic transducer surface, or an array of transducer surfaces, to generate an ultrasound beam. The transducer may be geometrically shaped and positioned to focus the ultrasonic energy at a “focal zone” corresponding to the target tissue mass within the patient. During wave propagation through the tissue, a portion of the ultrasound energy is absorbed, leading to increased temperature and, eventually, to cellular necrosis—preferably at the target tissue mass in the focal zone. The individual surfaces, or “elements,” of the transducer array are typically individually controllable, i.e., their phases and/or amplitudes can be set independently of one another (e.g., using a “beamformer” with suitable delay or phase shift in the case of continuous waves and amplifier circuitry for the elements), allowing the beam to be steered in a desired direction and focused at a desired distance and the focal zone properties to be shaped as needed. Thus, the focal zone can be rapidly displaced and/or reshaped by independently adjusting the amplitudes and phases of the electrical signal input into the transducer elements.
However, because the human body is flexible and moves even when a patient is positioned to keep still (due to respiration, for example, or small involuntary movements), treatment delivered as multiple sonications over time—even when delivered within seconds of each other—may require interim adjustments to targeting and/or to one or more treatment parameters. Compensation for motion is thus necessary to ensure that the ultrasound beam remains focused on the target and does not damage the surrounding healthy tissues.
Accordingly, an imaging modality, such as magnetic resonance imaging (MRI), may be used in conjunction with ultrasound focusing during non-invasive therapy to monitor the locations of both the target tissue and the ultrasound focus. Generally, an MRI system 100, as depicted in FIG. 1, includes a static-field magnet 102, one or more gradient-field coils 104, a radio-frequency (RF) transmitter 106, and an RF receiver (not shown). (In some embodiments, the same device is used alternately as RF transmitter or receiver.) The magnet includes a region 108 for receiving a patient 110 therein, and provides a static, relatively homogeneous magnetic field over the patient. Time-variable magnetic field gradients generated by the gradient-field coils 104 are superposed with the static magnetic field. The RF transmitter 106 transmits RF pulse sequences over the patient 110 to cause the patient's tissues to emit a (time-varying) RF response signal, which is integrated over the entire (two- or three-dimensional) imaging region and sampled by the RF receiver to produce a time series of response signals that constitute the raw image data. This raw data is passed on to a computation unit 112. Each data point in the time series can be interpreted as the value of the Fourier transform of the position-dependent local magnetization at a particular point in k-space (i.e., wavevector space), where the wavevector k is a function of the time development of the gradient fields. Thus, by Fourier-transforming the time series of the response signal, the computation unit 112 can reconstruct a real-space image of the tissue (i.e., an image showing the measured magnetization-affecting tissue properties as a function of spatial coordinates) from the raw data. The real-space magnetic-resonance (MR) image may then be displayed to the user. The MRI system 100 may be used to plan a medical procedure, as well as to monitor treatment progress during the procedure. For example, MRI may be used to image an anatomical region, locate the target tissue (e.g., a tumor) within the region, guide the beam generated by the ultrasound transducer 114 to the target tissue, and/or monitor the temperature in and surrounding the target tissue.
In image-guided systems, such as MRI-guided focused-ultrasound (MRgFUS) systems), motion compensation is generally accomplished by tracking the target (directly or indirectly) in the images and steering the ultrasound beam based on the tracked position. One approach to target tracking involves determining the coordinates of a set of one or more identifiable features, or “anatomical landmarks,” that can be located in each image; and computing the motion of the target, which is presumed to be at a known location relative to the landmarks, based on these coordinates. In an alternative approach, the relative shifts between successive images are determined by correlating one image with a large number of computationally shifted copies of the other image, and selecting the shifted image that provides the best match. In either case, significant image-processing time is expended to determine the target location, reducing the effective imaging rate and often impeding real-time motion compensation. In some cases, delays in recognizing and quantifying target motion cause beam-targeting inaccuracies within a tolerable range. Often, however, it becomes necessary to stop the treatment process and correct for any misalignment due to displacement of the target tissue or organ before treatment can be resumed. This results in significant inefficiencies in the treatment process, and may cause inconvenient delays.
Accordingly, there is a need for improved motion-tracking approaches that facilitate tracking the target, and compensating for its motion, in real time during treatment.